The present invention relates generally to magnetic resonance imaging (MRI) systems, and more particularly to an assembly designed to dissipate the heat generated by the gradient coils and RF coils that are used in MRI.
When a substance such as human tissue is subjected to a uniform magnetic field (polarizing field B0), the individual magnetic moments of the spins in the tissue attempt to align with the polarizing field, but precess about it in random order at their characteristic Larmor frequency. If the substance, or tissue, is subjected to a magnetic field (excitation field B1) which is in the x-y plane and which is near the Larmor frequency, the net aligned moment, or “longitudinal magnetization”, Mz may be rotated, or “tipped”, into the x-y plane to produce a net transverse magnetic movement Mt. A signal is emitted by the excited spins after the excitation signal B1 is terminated and this signal may be received and processed to form an image.
During patient scans, the gradient coils that produce the aforementioned magnet field dissipate large amounts of heat, typically in the order of tens of kilowatts. The majority of this heat is generated by resistive heating of the copper electrical conductors that form x, y, and z axis gradient coils when these coils are energized. The amount of heat generated is in direct proportion to the electrical power supplied to the gradient coils. The large power dissipation not only results in an increase in temperature to the gradient coil, the heat produced is distributed within the gradient coil assembly or resonance modules and influences the temperature in two other critical regions. These two regions are located at boundaries of the gradient assembly and include the patient bore surface and warm bore surface adjacent to the cryostat that houses the magnets. Each of these three regions has a specific maximum temperature limitation. In the resonance module, there are material temperature limitations such as the glass transition temperature. That is, although the copper and fiber reinforced backing of the coils can tolerate temperatures in excess of 120° C., the epoxy that is used to bond the layers together typically has a much lower maximum working temperature of approximately from 70° to 100° C. Regulatory limits mandate a peak temperature on the patient or surface of 41° C. The warm bore surface also has a maximum temperature that is limited to approximately 40° C. to prevent excessive heat transfer through the warm bore surface into the cryostat. Further, temperature variations of more than 20° C. can cause field homogeneity variations due to a temperature dependence of the field shim material that exhibits a magnetic property variation with temperature.
High current levels employed in conventional gradient coils produce significant heat proximate to the coil. This heat must be carried away from the coil and the magnet bore region to prevent damage to the coil and related structure, to avoid unwanted changes in the magnetic field due heating of magnet components, and to prevent unacceptable heating of a patient or other subject in the bore.
Cooling systems for gradient coils generally rely on conduction of the heat generated in the active circuits of the coil to water carrying pipes at some distance from the gradient coil, possibly as much as 10 mm away. The space between the active circuits and the water pipes is usually of material with good insulation properties, such as fiber-glass, making heat conduction inefficient. The water carrying pipes are also radially outward of the coil heat regions resulting in the hottest regions being nearest to the patient being scanned with no cooling directly between the hot regions and the patient. The resulting heat generation puts thermal limits on the operation of the coil. In general, increased peak strengths and high patient throughput are driving up operating currents and voltages. The increases in operating currents are generating additional heat loads surpassing the ability of existing thermal systems.
In general, prior devices have employed some form of coolant, usually water or ethylene glycol, and have provided thermal insulation. However, newer imaging protocols use higher power levels and further efforts are required to allow these advanced studies without exceeding temperature limits. Other devices have employed air cooling methods, with air being blown directly into the patient bore. The main limitation in this method is that patients frequently complain of being too cold. Yet another disadvantage is that the amount of air flow can vary significantly depending on the size of the patient so that in some cases insufficient or irregular flow will not cool the patient at all.